Implementation of colored wavelength shifters in phoswich detectors

ABSTRACT

A phoswich device for determining depth of interaction (DOI) includes a first scintillator having a first scintillation decay time characteristic, a second scintillator having a second scintillation decay time characteristic substantially equal to the first scintillation decay time, a photodetector coupled to the second scintillator, and a wavelength shifting layer coupled between the first scintillator and the second scintillator, wherein the wavelength shifting layer modifies the first scintillation decay time characteristic of the first scintillator to enable the photodetector to differentiate between the first decay time characteristic and the second decay time characteristic. The phoswich device is particularly applicable to positron emission tomography (PET) applications.

CROSS-REFERENCE TO RELATED APPLICATION AND CLAIM OF PRIORITY

This application is a non-provisional of and claims priority to under 35 U.S.C. §119(e) copending Application Ser. No. 61/100,916 filed Sep. 29, 2008.

Cross reference to related material can also be found in an application having U.S. application Ser. No.: 12/110,544, titled “Implementation of Wavelength Shifters in PHOSWICH Detectors”, which was filed on Apr. 28, 2008, the entire contents of which are incorporated herein by reference.

TECHNICAL FIELD

The present invention is in the field of nuclear medical imaging. In particular, the present invention relates to techniques for accurate detection of emission radiation in nuclear medical imaging processes such as positron emission tomography (PET).

BACKGROUND

Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.

There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a region of the body of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis.

In traditional PET imaging, a patient is injected with a radioactive substance with a short decay time. As the substance undergoes positron emission decay, it emits positrons which, when they collide with electrons in the patient's tissue emit two simultaneous gamma rays. The gamma rays emerge from the patient's body at substantially opposite directions. These rays eventually reach a scintillation device positioned around the patient. There is often a ring of scintillation devices surrounding the patient. When the gamma rays interact with oppositely positioned scintillation devices, light is emitted and detected. The light is usually transmitted through a lightguide to a photodetector. The light detected by the photodetector is then interpreted by a processor to enable an image of a slice of the region of interest to be reconstructed.

In PET (as well as SPECT) it is important to match the scintillator emission wavelength to the photodetector's optimal wavelength quantum efficiency (QE). For example, a typical photomultiplier tube (PMT) used in PET applications has a peak wavelength sensitivity at 420 nm while a typical LSO scintillator used in PET emits at 420 nm. Therefore, PMTs and LSO are very well matched in terms of wavelength matching. LSO is a very good scintillator for a PMT and is reasonably matched also for other silicon-based photodetectors such as avalanche photodiodes (APDs) and silicon photomultipliers (SiPMs). Scintillators for PET may be made from crystal materials such as, but not limited to, LSO, YSO, LYSO, LuAP (i.e., LuAlO₃:Ce), LuYAP, or LaBr3.

The phoswich approach has been used to improve the detection in PET applications by determining the depth-of-interaction (DOI) in the detector. PET scanners are typically made of long, thin detectors with high stopping power to meet high sensitivity requirements. In the absence of DOI information, however, the thickness of the scintillator reduces the spatial resolution due to parallax error. To compensate for reduced spatial resolution, detectors with DOI capability have been used. DOI capability can determine the location of the gamma interaction in the direction of the incident gamma (i.e., depth from the surface of the detector).

One way to implement DOI capability is to use a multi-layer detector, in which the layers are made of material with different scintillation properties. Because the layers have different characteristics, when a gamma event is detected it is possible to identify which layer absorbed the gamma photon and so to determine more accurately the spatial interaction location in three dimensions.

A conventional “phoswich” thus is a detector with two or more layers of different scintillators. Phoswich detectors comprising two or more scintillator layers offer a means to simultaneously achieve both high sensitivity and high spatial resolution in nuclear imaging. Each scintillator layer typically has a distinct decay time that allows the DOI of a gamma ray to be determined via pulse shape determination techniques. That is, layer identification is done by using differences in scintillation decay time inherent in the scintillators and pulse shape discrimination techniques.

The use of different types of scintillators in a phoswich may result in different light yields, emission spectra, densities, effective atomic numbers, and indices of refraction, which can often result in compromises in performance of the phoswich.

SUMMARY

The present invention solves the existing need in the art to determine the depth of interaction for a PET detector using the same or similar scintillation materials in a phoswich detector. An embodiment of the present invention uses a phoswich device for determining depth of interaction (DOI). The phoswich device includes a first scintillator having a first scintillation decay time characteristic and a second scintillator having a second scintillation decay time characteristic substantially equal to the first scintillation decay time. The phoswich device further includes a photodetector coupled to the second scintillator, and a wavelength shifting layer coupled between the first scintillator and the second scintillator, wherein the wavelength shifting layer modifies the first scintillation decay time characteristic of the first scintillator to enable the photodetector to differentiate between the first decay time characteristic (front scintillator) and the second decay time characteristic (back scintillator).

The phoswich device is particularly applicable to positron emission tomography (PET) applications.

BRIEF DESCRIPTION OF THE DRAWINGS:

The invention will now be described in greater detail in the following by way of example only and with reference to the attached drawings, in which:

FIG. 1 is a depiction of a conventional phoswich detector configuration.

FIG. 2 is a depiction of a phoswich detector where a wavelength shifting layer is sandwiched between two scintillators, in accordance with an embodiment of the present invention.

FIG. 3 a depiction of an exemplary phoswich detector where a wavelength shifting layer is sandwiched between two LSO scintillators, in accordance with an embodiment of the present invention.

FIG. 4 is a depiction of a PSD configuration, in accordance with a second embodiment of the present invention.

FIGS. 5A and 5B are graphs showing the pulse shape discrimination peaks (5A) obtained with the set-up in FIG. 4 and the energy spectra of the two scintillators (5B).

FIG. 6 is a schematic of a PET scanner using a phoswich device in accordance with an embodiment of the present invention.

DETAILED DESCRIPTION

As required, disclosures herein provide detailed embodiments of the present invention; however, the disclosed embodiments are merely exemplary of the invention that may be embodied in various and alternative forms. Therefore, there is no intent that specific structural and functional details should be limiting, but rather the intention is that they provide a basis for the claims and as a representative basis for teaching one skilled in the art to variously employ the present invention.

FIG. 1 depicts a conventional phoswich combination 100 that has been considered for Positron Emission Tomography (PET) imaging. The phoswich combination 100 has a first type scintillator crystal 110 made of LuAP. When a 511 keV gamma photon is absorbed in the LuAP crystal 110, it emits light scintillations around 360 nm, with a scintillation time of 30 ns. The light photons travel out of the LuAP scintillator array 110, and is absorbed by LSO scintillator array 120, and re-emitted at 420 nm with a scintillation time of 40 ns. The 420 nm enters into photodetector 130, where the light photons are converted into an electrical signal. An exemplary photodetector 130, such as a photomultiplier tube, detects the 420 nm light with decay time characteristics given as convolution between the decay time of the LuAP crystal and the decay time of the YSO crystal. The following equation shows how the decay time characteristics detected by the PMT 130 are obtained.

T(combined)=T(LuAP)

T(LSO)  Equation 1

The presence of T(combined) signifies an event in the front scintillator

As previously discussed, there may be problems with using different types of scintillator crystals in the same phoswich.

FIG. 2 is a depiction of a phoswich detector utilizing a wavelength shifting layer sandwiched between two scintillators, in accordance with an embodiment of the present invention. The phoswich device 200 includes a scintillator 202, a wavelength shifting layer 206, a filter 208, a scintillator 204 and a photodetecter 210. The photodetector can comprise a photomultiplier tube, a solid state photodetector such as an APD, MRS-PD and SiPM.

Wavelength shifting layer 206 may include, but is not limited to a plastic light guide, a crystal, and liquid coatings made from wavelength shifting material. An exemplary wavelength shifting material used was a green wavelength shifter. Specifically, an Eljen 280 foil (WLS) that has an excitation band around 430 nm, suitable to interact with light emission from LSO, and an emission band around 500 nm. The WLS foil has a thickness around 0.1 mm.

It should be appreciated by those skilled in the art that although a green wavelength shifter is used, the invention may be modified to use other wave length shifting layers as long as interactions occur between the front scintillator and the shifting layer and a decay time modification takes place. Other scintillator types may fall within the scope of the present invention.

In a first embodiment of the invention, scintillator 202 and scintillator 204 are from the same crystal group which results in scintillator 202 and scintillator 204 having substantially equal decay time characteristics. For example, scintillator 202 comprises a LSO crystal and scintillator 204 comprises a LSO crystal. Scintillator 202 and scintillator 204 may comprise LSO, YSO, LUAP, LUYap, LFS, LYSO, LaBr3, and the like crystals.

In a second embodiment of the invention, scintillator 202 and scintillator 204 are from different crystal groups but both crystals for light transmission purposes are substantially alike, e.g. have substantially the same decay time characteristics and do not interact.

In an embodiment of the invention, a 511 keV gamma photon is absorbed by scintillator 202. The gamma photon emits light scintillations around 420 nm, with a scintillation time of 40 ns. The light photons exit scintillator 202 and are absorbed by wavelength shifting material 206. Wavelength shifting material 206 can affect the light photons in a number of ways depending on the type of wavelength absorbing material used. For instance, in one embodiment of the invention, the wavelength shifting material 206 increases the decay time characteristic of the light photons entering scintillator 204. In another embodiment of the invention, the decay time remains substantially the same, however, the rise time of the light photons entering scintillator 204 is affected.

The light photons exit the wavelength shifting material 206 and enter scintillator 204. The light photons then exit scintillator 204 with a decay time characteristic of 40 ns but a modified rise time and enter the photodetector 210 where the light photon is converted to an electrical signal.

The filter 208, which is preferably a long pass filter prevents light from scintillator 204 from being reflected into scintillator 202.

Conversely, a 511 keV photon travels through scintillator 202, and is absorbed by the scintillator 204. Scintillator 204 now emits 420 nm light with a decay time of 40 ns. Thus, identification of the location of the gamma interaction in either the front scintillator or the back scintillator can be easily made by analyzing the signals from photodetector 210.

FIG. 3 a depiction of an exemplary phoswich detector 300 where a wavelength shifting layer 306 is sandwiched between two LSO scintillators 302, 304 in accordance with an embodiment of the present invention. In an embodiment of the invention, a 511 keV gamma photon is absorbed by LSO scintillator 302. The gamma photon emits light scintillations of 420 nm, with a scintillation time of 40 ns. The light photons exit LSO scintillator 302 and are absorbed by wavelength shifting material 306. The LSO decay time characteristics is changed due to the interaction and the light emission is shifted upwards to around 500 nm. The shifted light has, thus, the decay time characteristics of a convolution between LSO light response and the response of the wavelength shifter. Specifically, the rise time of the light signal is affected by wavelength shifting layer 306. Specifically, the light signal has a 12 ns rise time. The decay time remains substantially the same.

The light photons generated in 302 exit the wavelength shifting material 306 and enter the photodetector 310 via the LSO scintillator 304.

Conversely, a 511 keV photon travels through LSO scintillator 302, and is absorbed in LSO scintillator 304 which emits 420 nm light with a decay time of 40 ns which is detected by photodetector 310. Filter 308, prevents light from LSO scintillator 304 from being transmitted in the wave length shifter and into LSO scintillator 302. One light signal arrives with a light scintillation of 500 nm and a rise time of 12 ns and a 40 ns decay time signifying an event in the front scintillators (302), another signal arrives with a light scintillation of 420 nm and a rise time of 1 ns and a 40 ns decay time, signifying a 511 keV event has been registered in the back scintillator (304). With the rise time sensitive pulse shape discriminator circuit, the two light signals can be differentiated.

FIG. 4 is a depiction of a Pulse Shape Discrimination (PSD) circuit 400 in accordance with a second embodiment of the present invention. Signal 402 also known as signal A comprises an exemplary 500 nm light scintillation with a 12 ns rise time. Signal 404 also known as signal B comprises a 420 nm light scintillation with a 1 ns rise time. Signal A and signal B are detected by a photodetector 406 depicted as a PMT. Two constant fraction discriminator (CFD) circuits are provided. The anode signal goes to discriminator circuit 410, timing circuit 412 and discriminator circuit 414 and time activity curve circuit (TAC) 416. The result is illustrated in graph 420 which depicts the time channel of signal A and Signal B. The dynode signal if available is connected to a spectroscopy amplifier to provide the phoswich energy spectra. Graph 418 illustrates the energy channel of signal A and signal B

FIGS. 5A and 5B are graphs illustrating a time spectrum and an energy spectrum for the phoswich of FIG. 3 in accordance with an embodiment of the present invention. For FIG. 5A, which depicts the PSD time spectrum all events below channel 88 are defined as fast events (e.g. scintillator 304 being closest to the photodetector 310). All events greater than channel 88 are defined as slow events (longer rise time, scintillator 302). Based on these two gates, the two energy spectra depicted in FIG. 5B are acquired. The source ⁶⁸Ge, providing the 511 keV photons, was positioned just above scintillator 302. Based on the two Gaussians from FIG. 5A, the cross-talk between scintillator 302 and scintillator 304 can be calculated. For the fast setting (e.g., channels below 88) there are 93% true fast events and 7% slow events coming in via scintillator 304. For the slow setting, the numbers are 98% true slow events and 2% fast events.

FIG. 6 is a diagram of a PET scanning system 600 using a wavelength shifting material in the phoswich device in accordance with another embodiment of the invention. PET scanning system 600 consists of a number of phoswich detectors 620. The phoswich detectors may be arranged in a ring configuration. The ring of phoswich detectors 620 forms a space large enough for an adult human body to pass. Each phoswich detector may consist of a first scintillator material, a wavelength shifting material, a second scintillator material and a photodetector. The ring of phoswich detectors 620 may be connected to a processor 630. The processor 630 is capable of analyzing the data received from the ring of phoswich detectors 620, reconstructing an image from the acquired data, and outputting tomographic images of the object or patient scanned. The PET scanning system 600 may further include a table or other support structure 610 capable of holding the object or patient to be scanned. The table or other support structure 610 may be adapted to pass through the bore formed by the ring of block detectors 620.

The invention having been thus described, it will be apparent to those skilled in the art that the same may be varied in many ways without departing from the spirit and scope of the invention. Any and all such modifications are intended to be covered within the scope of the following claims. 

1. A phoswich detector comprising: a first scintillator having a first scintillation decay time characteristic; a second scintillator having a second scintillation decay time characteristic substantially equal to the first scintillation decay time; a photodetector coupled to the second scintillator; and a wavelength shifting layer coupled between the first scintillator and the second scintillator, wherein the wavelength shifting layer modifies a first scintillation rise time characteristic of the first scintillator to enable the photodetector to differentiate between the scintillation characteristics of a front and a back scintillator.
 2. The phoswich detector of claim 1, further comprising: a long pass filter for preventing light from the second scintillator from entering the first scintillator.
 3. The phoswich detector of claim 1, further comprising: a pulse shape discriminator circuit for differentiating between a pulse shape associated with the first scintillator and a pulse shape associated with the second scintillator.
 4. The phoswich detector of claim 1, wherein the wavelength shifting layer comprises a green wavelength shifter.
 5. The phoswich detector of claim 1, wherein the wavelength shifting layer comprises at least one of a plastic light guide, a crystal, and a liquid coating.
 6. The phoswich detector of claim 1, wherein the photodetector comprises a photomultiplier tube.
 7. The phoswich detector of claim 1, wherein the photodetector comprises a solid state photodetector.
 8. The phoswich detector of claim 7, wherein the photodetector comprises at least one of APDs, MRS-PDs and SiPMs.
 9. The phoswich detector of claim 1, wherein the first scintillator and the second scintillator comprise LSO crystals.
 10. The phoswich detector of claim 1, wherein the first scintillator and the second scintillator comprise LaBr3 crystals.
 11. The phoswich detector of claim 1, wherein the first scintillator and the second scintillator comprise YSO crystals.
 12. The phoswich detector of claim 1, wherein the first scintillator and the second scintillator comprise LuAP crystals.
 13. The phoswich detector of claim 1, wherein the first scintillator and the second scintillator comprise LuYap crystals.
 14. The phoswich detector of claim 1, wherein the first scintillator and the second scintillator are both selected from the group consisting of LSO, LaBr3, YSO, LUAP, LUYap, LYSO, and LFS.
 15. A medical imaging apparatus comprising: a plurality of scintillation detector modules, each module comprising a first scintillator having a first scintillation decay time characteristic; a second scintillator having a second scintillation decay time characterstic substantially equal to the first scintillation decay time; a photodetector coupled to the second scintillator; a wavelength shifting layer coupled between the first scintillator and the second scintillator, wherein the wavelength shifting layer modifies the first scintillation decay time characteristic of the first scintillator to enable the photodetector to differentiate between the first decay time characteristic and the second decay time characteristic; a processor for receiving medical imaging data from the detector modules; and software executing on the processor for analyzing medical imaging data from the detector modules and for reconstructing a tomographic image based on the medical imaging data.
 16. The medical imaging scanner of claim 15, wherein the medical imaging scanner comprises a Positron Emission Tomography (PET) scanner.
 17. The medical imaging scanner of claim 15, wherein the medical imaging data comprises a Positron Emission Tomography (PET) data.
 18. The medical imaging scanner of claim 15, wherein each module further comprises a long pass filter for preventing light from the second scintillator from entering the first scintillator.
 19. The medical imaging scanner of claim 15, wherein each module further comprises a pulse shape discriminator circuit for differentiating between a pulse shape associated with the first scintillator and a pulse shape associated with the second scintillator.
 20. The medical imaging scanner of claim 15, wherein the wavelength shifting layer comprises a green wavelength shifter. 